Determination of leak during CPAP treatment

ABSTRACT

A respiratory treatment apparatus and method in which a leak is determined by using an averaging window. The window starts at the present time and extends back in time to a point determined according to a current one of progressively detected phase measures of a first respiratory cycle and a corresponding phase measure attributable to a preceding second respiratory cycle. In another aspect, a jamming index indicates whether the leak is rapidly changing. To the extent that jamming is high, the leak estimate used progressively changes from that using sliding breath-window averaging to a more robust and faster responding low-pass filter method, and adjustment of ventilatory support based on measures employing estimated respiratory flow is slowed down or stopped.

CROSS-REFERENCE TO RELATED APPLICATIONS

The present application is a divisional of U.S. patent application Ser.No. 13/874,668, filed May 1, 2013, which is a continuation of U.S.patent application Ser. No. 13/542,331, filed Jul. 5, 2012, now U.S.Pat. No. 8,459,260, which is a continuation of U.S. patent applicationSer. No. 12/438,758, filed Feb. 25, 2009, now U.S. Pat. No. 8,256,418which is a national stage application of PCT Application No. PCT/AU2007/001237, filed Aug. 30, 2007, which claims the benefit of the filingdate of U.S. Provisional Application 60/823,934, filed Aug. 30, 2006,the disclosures of which are incorporated herein by reference.

FIELD OF THE INVENTION

This invention relates to treatment of apneas and other respiratorydisorders. In particular it relates to methods and apparatus for thedetermination of leakage airflow and true respiratory airflow, duringthe mechanical application of positive airway pressure.

BACKGROUND OF THE INVENTION

For the treatment of apneas and other respiratory disorders, breathablegas is supplied from a mechanical respirator or ventilator, for examplevia a mask, at a pressure which may be higher during inspiration andlower during expiration. (In this specification any reference to a“mask” is to be understood as including all forms of devices for passingbreathable gas to a person's airway, including nose masks, nose andmouth masks, nasal prongs/pillows and endrotracheal or tracheostomytubes. The term “ventilator” is used to describe any device that doespart of the work of breathing.) Typically one measures the subject'srespiratory airflow during mechanical ventilation to assess adequacy oftreatment, or to control the operation of the ventilator.

Respiratory airflow is commonly measured with a pneumotachograph placedin the gas delivery path between the mask and the pressure source. Leaksbetween the mask and the subject are unavoidable. The pneumotachographmeasures the sum of the respiratory airflow plus the flow through theleak plus flow through the vent (also called “deliberate leak”). If theinstantaneous flow through the leak is known, the respiratory airflowcan be calculated by subtracting the flow through the leak and the flowthrough the vent from the flow at the pneumotach. Typically the flowthrough the vent is a known function of pressure at the vent, and giventhat the pressure at the vent can be estimated with reasonable accuracy,the flow through the vent can then be straightforwardly calculated. Onthe other hand, if the vent characteristics are suitable for the leakmodel employed, the vent flow and non-deliberate leak can be lumpedtogether and estimated as a single quantity. The direct estimation ofvent flow using pressure at the vent will be assumed hereinafter, andsubtraction of this vent flow from total ventilator outflow will beassumed to have occurred when not mentioned explicitly.

Some methods to correct for the flow through the leak assume (i) thatthe leak is substantially constant, and (ii) that over a sufficientlylong time, inspiratory and expiratory respiratory airflow will cancel.If these assumptions are met, the average flow through the pneumotachover a sufficiently long period will equal the magnitude of the leak,and the true respiratory airflow may then be calculated as described.

The known method is only correct if the pressure at the mask isconstant. If, on the other hand, the mask pressure varies with time (forexample, in the case of a ventilator), assumption (i) above will beinvalid, and the calculated respiratory airflow will therefore beincorrect. This is shown markedly in FIGS. 1a -1 f.

FIG. 1a shows a trace of measured mask pressure in bi-level CPAP(Continuous Positive Airway Pressure) treatment between about 4 cm H2Oon expiration and 12 cm H2O on inspiration. FIG. 1b shows a trace oftrue respiratory airflow in synchronism with the mask pressures. Attime=21 seconds a mask leak occurs, resulting in a leakage flow from theleak that is a function of the treatment pressure, as shown in FIG. 1 c.The measured mask flow shown in FIG. 1d now includes an offset due tothe leak flow. The prior art method then determines the calculated leakflow over a number of breaths, as shown in FIG. 1 e. The resultingcalculated respiratory flow, as the measured flow minus the calculatingleak flow is shown in FIG. 1 f, having returned to the correct meanvalue, however is incorrectly scaled in magnitude, giving a falseindication of peak positive and negative airflow.

Another prior art arrangement is disclosed in European Publication No. 0714 670 A2, including a calculation of a pressure-dependent leakcomponent. The methodology relies on knowing precisely the occurrence ofthe start of an inspiratory event and the start of the next inspiratoryevent. In other words, the leak calculation is formed as an average overa known breath and applied to a subsequent breath.

This method cannot be used if the moment of start and end of theprevious breath are unknown. In general, it can be difficult toaccurately calculate the time of start of a breath. This is particularlythe case immediately following a sudden change in the leak.

Furthermore, the method will not work in the case of a subject who ismaking no respiratory efforts, and is momentarily not being ventilatedat all, for example during an apnea, because for the duration of theapnea there is no start or end of breath over which to make acalculation.

In U.S. Pat. No. 6,152,129 (Berthon-Jones) the leak is determined byfirst estimating the conductance of the leak path from the long termorifice flow:

${\frac{1}{R_{L}} = \frac{< Q >}{< \sqrt{p} >}},$where G_(L)=1/R_(L) is conductance (L denotes leak), Q is instantaneousflow, p is instantaneous pressure and < >denotes a long term averagecalculated for example by low pass filtering with an IIF or other filterhaving a long time constant. Note that the word “average” as used hereincontains the general sense inclusive of the result of a low passfiltering step, and is not limited to an arithmetic mean or otherstandard average such as the RMS average.The instantaneous leak flow, based on the model of the flow through anorifice is then

$Q_{L} = {\frac{1}{R_{L}}\left. \sqrt{}p \right.}$

Note that the instantaneous respiratory airflow is thenQ _(R) =Q−Q _(L).

Berthon-Jones attempts to deal with sudden changes in instantaneous leakflow by dynamically adjusting the filter's time constant using fuzzylogic, lengthening the time constant if it is certain that the leak issteady, reducing the time constant if it is certain that the leak hassuddenly changed, and using intermediately longer or shorter timeconstants if it is intermediately certain that the leak is steady.

Berthon-Jones also develops a jamming index by fuzzy logic to deal withthe case of a large and sudden increase in the conductance of the leak,in which case the calculated respiratory airflow will be incorrect. Inparticular during apparent inspiration, the calculated respiratoryairflow will be large positive for a time that is large compared withthe expected duration of a normal inspiration. Conversely, if there is asudden decrease in conductance of the leak, then during apparentexpiration the calculated respiratory airflow will be large negative fora time that is large compared with the duration of normal expiration.

Therefore, the jamming index, i.e. an index of the degree of certaintythat the leak has suddenly changed, is derived, such that the longer theairflow has been away from zero, and by a larger amount, the larger theindex. The explicit calculation of the jamming index by fuzzy logic isdescribed in the '129 patent, which is incorporated herein by reference.

The time constant for the low pass filters is then adjusted to varyinversely with the jamming index. In operation, if there is a sudden andlarge change in the leak, the index will be large, and the time constantfor the calculation of the conductance of the leak will be small,allowing rapid convergence on the new value of the leakage conductance.Conversely, if the leak is steady for a long time, the index will besmall, and the time constant for calculation of the leakage conductancewill be large; enabling accurate calculation of the instantaneousrespiratory airflow. In the spectrum of intermediate situations, wherethe calculated instantaneous respiratory airflow is larger and forlonger periods, the index will be progressively larger, and the timeconstant for the calculation of the leak will progressively reduce. Forexample, at a moment in time where it is uncertain whether the leak isin fact constant, and the subject merely commenced a large sigh, orwhether in fact there has been a sudden increase in the leak, the indexwill be of an intermediate value, and the time constant for calculationof the impedance of the leak will also be of an intermediate value.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1a shows a trace of measured marked pressure in bi-level CPAPtreatment.

FIG. 1b shows a trace of true respiratory airflow in synchronism withmark pressures.

FIG. 1c shows a trace of leakage flow that is a function of treatmentpressure.

FIG. 1d shows a trace that illustrates an offset due to leak flow.

FIG. 1e shows calculated leak flows over a number of breaths.

FIG. 1f shows calculated respiratory flow.

FIG. 2 illustrates elements of a system 200 of the present technology.

BRIEF DESCRIPTION OF THE INVENTION

This invention rapidly determines the instantaneous leak in a CPAPsystem without detailed modeling the source of the leak and withouthaving to determine the precise phase in a breathing cycle at which theleak occurs. It relies instead on the use of timers to define thebreathing cycle and a calculation to assure that the instantaneous flowis compared to the flow over a time period long enough to include anentire breath. It does this by looking backward to include an entirephase cycle. This avoids having to take long term averages over multiplebreaths, or to have a model that recognized the beginning and end of abreath.

Sudden changes in a leak are recognized and the degree to which leak israpidly changing is expressed as a jamming index value, which is thenused as a parameter to adjust the contributions of the components ofwhich the leak estimate is made up, and, in the case of aservoventilator, to temporarily slow down or suspend the adjustment ofthe servoventilator controller output parameter, typically pressuresupport level.

DETAILED DESCRIPTION OF PREFERRED EMBODIMENTS

The present invention is motivated by the desire to produce acontinuously updated estimate of the leak model parameter which is verystable when the actual leak parameter is stable, but which degradesprogressively and gracefully to a less stable, less accurate butfaster-responding estimate when the actual leak parameter is changingrapidly. The leak model parameter in question is typically an orificeconstant (equivalently a leak conductance), but need not be.

A continuously updated estimate of leak conductance (in particular,continuously updated during each breath) may be calculated by performingsome kind of low-pass filtering operation, such as a 1st-order low passfilter or a moving average filter, typically with a fixed window length,on the non-vent flow (equal to the sum of the respiratory flow and theinstantaneous leak flow) and on the square root of the mask pressure,producing a leak conductance estimate G₁, as in Berthon-Jones. Thismethod has the advantage over some other methods that it is independentof the determination of breath phase (the position within the currentbreath). Thus sudden changes in leak can occur which cause respiratoryflow estimates and hence breath phase estimates to be grossly in error,yet updating of the leak parameter estimates continues. A disadvantageof such breath-independent estimates is that the estimates are notstable within a breath unless particular fortuitous events occur; forexample, by coincidence, the duration of a window averaging filterincludes exactly N breaths, where N is an integer. A particular case ofthis instability is illustrated by considering the situation when 1storder low-pass-filter estimates of mask flow and mask pressure are used.For simplicity, assume that a mask pressure is constant, and that trueleak conductance is zero. Then the leak flow estimate is just a1st-order low-pass-filtered version of respiratory flow. This estimaterises whenever respiratory flow is above the leak flow estimate, andfalls when respiratory flow is below the leak flow estimate. Inparticular, with reasonable filter time constants, the leak flowestimate rises during most of inspiration and falls during most ofinspiration, rising slowly during the expiratory pause, and under normalcircumstances crucially being below zero during the expiratory pause.Since true respiratory flow is zero during the expiratory pause,estimated respiratory flow, being the difference between mask flow (byassumption equal to respiratory flow) and estimated leak flow, ispositive during the expiratory pause, say equal to Q_(eps). If aventilator is designed to trigger into inspiration when the estimatedrespiratory flow exceeds some true respiratory flow thresholdQ_(insp_thresh), a ventilator which uses this kind of leak estimate, inorder to trigger at the desired true respiratory flow, must set itstrigger threshold to a higher value Q_(insp_thresh)+Q_(eps).

Unfortunately Q_(eps) is a function of the respiratory flow shape, therespiratory rate, and the low-pass-filter time constant, very difficultif not impossible to determine in real time. Hence triggering actuallyoccurs at a variable threshold, and in the worst case auto-triggering(triggering at a true respiratory flow of zero) may occur. It should benoted that the effect of the non-constant estimate of leak parameter inproducing a distorted respiratory flow exists throughout the breath andwhether there is an identifiable expiratory pause or not, with potentialadverse effects on cycling (expiratory triggering) as well an on(inspiratory) triggering, as well as other algorithms which operate onestimated respiratory flow.

“Jamming”, as described by Berthon-Jones, is the extent to which theleak has not yet been compensated for, and usually results from a rapidchange in the leak. It is herein considered to be a fuzzy logicalquantity.

A leak conductance estimate G₁ is calculated as described above. Notethat the time constant of the filters uses preferably decreases asjamming increases, as described in Berthon-Jones.

A second leak conductance estimate G₂ is calculated continuously, at thealgorithmic sampling frequency or a lower frequency (e.g. 10 Hz) whichis still high compared with the respiratory frequency. In a mannerdescribed below, the algorithm identifies the position in the currentbreath, then attempts to find time associated with the same position inthe preceding breath. If it fails to find such a position, it usesinstead a time 10 seconds in the past. Between that time in the past andthe present, a window is established. Low-pass filtered mask flow andlow-pass filtered square root of mask pressure (filtered by anon-breath-dependent method, such as a 1st-order LPF), typically thelow-pass filtered values used for the determination of G₁, are thenfurther low-pass filtered by being averaged over this window. The ratioof these window-averaged values is the leak conductance estimate G₂,which under conditions of stable leak is extremely stable.

Because G₂ responds rather slowly to changes in leak conductance, it isinappropriate to use when the leak is changing rapidly. Thus to theextent that there is jamming, G₁ rather than G₂ is used. In thepreferred implementation, if J is jamming (a quantity in [0,1]), theconductance estimate G_(j) is used, given byG _(j) =JG ₁+(1−J)G ₂Instantaneous leak is then straightforwardly calculated byQ _(leak) =G _(j)√{square root over (P _(mask))}.

Determining the Position of the Start of the Averaging Window for G₂

The aim is to determine the same position in the previous breath as thepatient, is at in the current breath. For this one needs a concept ofbreath phase which is not just one of a small set of categories, such asinspiration and expiration, but a conceptually real-valued (in practicerational) variable which increases progressively from the start ofinspiration to the end of expiration, potentially with a small finitenumber of jumps. Such a concept is provided in Berthon-JonesCheyne-Stokes patent, WO98/012965, which is incorporated herein byreference. Breath phase is there defined to be 0 at the start ofinspiration, 0.5 at the start of expiration, and approaches 1 at the endof expiration. Given such a breath phase, one find the breath phase atthe current moment, then searches backward in time to find the samebreath phase in the previous breath. Because breath phase as estimatedby the system described by Berthon-Jones is not necessarily increasingwith time during a breath (neglecting the expiratory to inspiratorytransition, at which it must decrease) though typically it is increasingwith time during a breath, it is necessary to have an algorithm whichsearches backward in time in such a way that a point in the same breathwith the same breath phase as the current value is not identified asbeing in the previous breath. Such a search algorithm is describedbelow; this algorithm may fail under exceptional circumstances, but isquite robust most of the time. Because of jumps in phase, there mayexist no point in the previous breath with a phase equal to the phaseassociated with the current moment, the latest time in the previousbreath with a phase less than or equal to the phase at the currentmoment is used instead.

On the other hand, a system which uses conventional flow thresholds fortriggering and cycling need not use a fuzzy logical system fordetermining breath phase for the purpose of finding the same position inthe previous breath as in the current breath. Assuming that duringinspiration, the maximum time between the present until the end ofinspiration is known (typically determined at the start of inspiration,but not necessarily), the breath phase at each sampling interval isincreased by such an amount that with equal increments of that amount,the phase would reach 0.5 at the end of inspiration. For example, in thesimple case where a maximum inspiratory time of 1.6 seconds wasdetermined at the start of inspiration, the phase would increase at asteady rate of 0.5/1.6 phase units/second. If cycling (transition toexpiration) occurred earlier, for example because respiratory flow fellbelow a cycling threshold, the phase would at that point jump to 0.5.Similar considerations apply during expiration, with rate of increase ofphase being the difference between 1 and the current phase divided bythe time remaining until the maximum expiratory time. If since breathphase determined in this way is typically used only for the purpose ofdetermining the same position in the previous breath as in the currentbreath, it is called “book-keeping” phase.

Regardless of the phase determination method used, whether that ofBerthon-Jones, “book-keeping” phase as described above, or some othermethod, the search backward in time to find the latest time in thepreceding breath with a phase less than or equal to the phase at thecurrent moment is performed as follows (though it will be appreciatedthat for “book-keeping” phase, simpler methods are available).

Starting with the current phase, say φ₀, the invention looks backwardsin time for the most recent phase in the interval [φ₀−0.75, φ₀−0.25].The aim is to seek a point in time at least 0.25 of a breath before thepresent. When such a phase is found, the invention calculates φ₁=φ₀−0.25and looks backward for a phase in the interval [φ₁−0.75, φ₁−0.25]. Thisis continued, 0.25 at a time, i.e. φ_(i+1)=φ_(i)−0.25. When a phase isfound which is in [φ₃−0.075, φ₃−0.25] the iteration ceases, since thisis just [φ₀−0.5, φ₀]. If phase varied continuously this would have foundexactly φ₀; in reality it will most likely find φ₀−ε., where hopefully εis small. By proceeding in this manner we have some confidence that thephase has gone backward rather than forward, since we have found phasesin the 4 phase quadrants before the present. This algorithm will regardtwo phase transitions of 0.5 in succession as being movement backward,though the actual direction is of course actually indeterminate. If thisalgorithm fails to find a point between the present moment and a timebefore the present which meets this criterion, we take the start of theaveraging window to be some reasonable maximum time before the present,such as 10 seconds. As an implementation detail, to reduce computationalrequirements, the leak, flow values may be averaged over the last 0.1seconds and stored in a buffer accompanied by the associated breathphase, so that the search for the last breath is performed in a bufferof 100 points, and done every 0.1 seconds. The averaged leak estimate atthe instantaneous leak calculation frequency, e.g. 100 Hz, can then becalculated by linear interpolation between the most recent averaged leakconductance estimate and the averaged conductance leak estimate justbefore it.

In a servoventilator or other system using some kind of measure ofventilation (such as half the absolute value of respiratory flow, or agross alveolar ventilation, or peak flow, or some weighted average offlows determined during inspiration or expiration) to adjust ventilatorysupport, when jamming is observed, the system slams down or suspendschanges in pressure support. This is because respiratory flow estimatesare not reliable in the presence of jamming, and various measures ofventilation based on respiratory flow are likely to overestimateventilation, leading for example in a servoventilator to unwarrantedwithdrawal in ventilatory support because ventilation appears to beabove target ventilation. The extent of slowing down of adjustment ofrespiratory support is preferably some increasing function of thejamming. For example, if the calculated change in respiratory supportfrom that at the previous time that it was calculated is some value ΔS,then the adjusted change in support would be kΔS, where for example k is1 for J≤0.1, 0 for J≥0.3, and taking linearly interpolated values forintermediate values of J.

The invention claimed is:
 1. An apparatus for providing pressurizedbreathable gas to a subject, comprising: a source for generating thepressurized breathable gas; a sensor for measuring respiratory flowassociated with ventilation of the subject; and a controller thatoutputs an output parameter associated with a pressure support level ofthe breathable gas, wherein the pressure support level is adjusted basedon an index value associated with non-deliberate leak of the pressurizedbreathable gas, wherein the controller is configured to adjust thepressure support level according to a result of a multiplication of acalculated pressure support change value and an increasing function ofthe index value.
 2. The apparatus according to claim 1, wherein theindex value is less than
 1. 3. The apparatus according to claim 1,wherein if the index value is less than or equal to 0.1, the pressuresupport level is adjusted.
 4. The apparatus according to claim 3,wherein the pressure supported level is adjusted to be a change inpressure support from a previous time.
 5. The apparatus according toclaim 1, wherein the index value is calculated using fuzzy logic.
 6. Theapparatus according to claim 1, wherein the apparatus comprises aservoventilator and the controller comprises a servoventilatorcontroller.
 7. The apparatus according to claim 6, wherein theservoventilator controller reduces changes in the pressure support levelfor a predetermined time period.
 8. The apparatus according to claim 6,wherein the servoventilator controller stops changes in the pressuresupport level for a predetermined time period.
 9. The apparatusaccording to claim 1, wherein the apparatus comprises a mechanicalventilator.
 10. The apparatus according to claim 1, wherein theincreasing function of the index value comprises linearly interpolatedvalues.
 11. An apparatus that uses a measure of ventilation to adjust aventilatory support, comprising: a sensor for measuring ventilationassociated with a subject; and a controller having an output parameter,the output parameter being associated with an index value which providesan indication of a degree to which leak associated with the measure ofventilation associated with the subject is changing, wherein thecontroller is configured to adjust a pressure support level of theventilatory support according to a result of a multiplication of acalculated pressure support change value and an increasing function ofthe index value.
 12. The apparatus according to claim 11, wherein theapparatus further comprises a mask that delivers pressurized breathablegas to the subject and the index value represents a jamming index ofuncompensated leak of the mask.
 13. The apparatus according to claim 11,wherein if the index value is less than or equal to 0.1, the pressuresupport level is adjusted.
 14. The apparatus according to claim 13,wherein the pressure supported level is adjusted to be a change inpressure support from a previous time.
 15. The apparatus according toclaim 11, wherein if the index value is equal to or greater than 0.3,the pressure support level is not adjusted.
 16. The apparatus accordingto claim 11, wherein if the index value is greater than 0.1 and lessthan 0.3, the pressure support level is adjusted based linearlyinterpolated values of the index value.
 17. The apparatus of claim 11,wherein the pressure support level is adjusted so as to slow downchanges associated with the pressure support level.
 18. The apparatus ofclaim 11, wherein the pressure support level is adjusted so as tosuspend changes associated with the pressure support level.
 19. Theapparatus according to claim 11, wherein the apparatus comprises aservoventilator and the controller comprises a servoventilatorcontroller.
 20. A method for adjusting a level of ventilatory support ofa ventilator, comprising: measuring respiratory flow associated withventilation of a subject; deriving a jamming index value which providesa measure of change in uncompensated leak; and adjusting a pressuresupport level of the ventilatory support based on the jamming index suchthat adjustment of the ventilatory support is applied according to aresult of a multiplication of a calculated pressure support change valueand an increasing function of the jamming index.
 21. The methodaccording to claim 20, wherein if the jamming index value is less thanor equal to 0.1, the pressure support level is adjusted.